1. Field of the Invention
The present invention relates to tissue-supporting medical devices, and more particularly to expandable, non-removable devices that are implanted within a bodily lumen of a living animal or human to support the organ and maintain patency.
2. Summary of the Related Art
In the past, permanent or biodegradable devices have been developed for implantation within a body passageway to maintain patency of the passageway. These devices are typically introduced percutaneously, and transported transluminally until positioned at a desired location. These devices are then expanded either mechanically, such as by the expansion of a mandrel or balloon positioned inside the device, or expand themselves by releasing stored energy upon actuation within the body. Once expanded within the lumen, these devices, called stents, become encapsulated within the body tissue and remain a permanent implant.
Known stent designs include monofilament wire coil stents (U.S. Pat. No. 4,969,458); welded metal cages (U.S. Pat. Nos. 4,733,665 and 4,776,337); and, most prominently, thin-walled metal cylinders with axial slots formed around the circumference (U.S. Pat. Nos. 4,733,665, 4,739,762, and 4,776,337). Known construction materials for use in stents include polymers, organic fabrics and biocompatible metals, such as, stainless steel, gold, silver, tantalum, titanium, and shape memory alloys such as Nitinol.
U.S. Pat. Nos. 4,733,665, 4,739,762, and 4,776,337 disclose expandable and deformable interluminal vascular grafts in the form of thin-walled tubular members with axial slots allowing the members to be expanded radially outwardly into contact with a body passageway. After insertion, the tubular members are mechanically expanded beyond their elastic limit and thus permanently fixed within the body. The force required to expand these tubular stents is proportional to the thickness of the wall material in a radial direction. To keep expansion forces within acceptable levels for use within the body (e.g., 5-10 atm), these designs must use very thin-walled materials (e.g., stainless steel tubing with 0.0025 inch thick walls). However, materials this thin are not visible on conventional fluoroscopic and x-ray equipment and it is therefore difficult to place the stents accurately or to find and retrieve stents that subsequently become dislodged and lost in the circulatory system.
Further, many of these thin-walled tubular stent designs employ networks of long, slender struts whose width in a circumferential direction is two or more times greater than their thickness in a radial direction. When expanded, these struts are frequently unstable, that is, they display a tendency to buckle, with individual struts twisting out of plane. Excessive protrusion of these twisted struts into the bloodstream has been observed to increase turbulence, and thus encourage thrombosis. Additional procedures have often been required to attempt to correct this problem of buckled struts. For example, after initial stent implantation is determined to have caused buckling of struts, a second, high-pressure balloon (e.g., 12 to 18 atm) would be used to attempt to drive the twisted struts further into the lumen wall. These secondary procedures can be dangerous to the patient due to the risk of collateral damage to the lumen wall.
Many of the known stents display a large elastic recovery, known in the field as “recoil,” after expansion inside a lumen. Large recoil necessitates over-expansion of the stent during implantation to achieve the desired final diameter. Over-expansion is potentially destructive to the lumen tissue. Known stents of the type described above experience recoil of up to about 6 to 12% from maximum expansion.
Large recoil also makes it very difficult to securely crimp most known stents onto delivery catheter balloons. As a result, slippage of stents on balloons during interlumenal transportation, final positioning, and implantation has been an ongoing problem. Many ancillary stent securing devices and techniques have been advanced to attempt to compensate for this basic design problem. Some of the stent securing devices include collars and sleeves used to secure the stent onto the balloon.
Another problem with known stent designs is non-uniformity in the geometry of the expanded stent. Non-uniform expansion can lead to non-uniform coverage of the lumen wall creating gaps in coverage and inadequate lumen support. Further, over expansion in some regions or cells of the stent can lead to excessive material strain and even failure of stent features. This problem is potentially worse in low expansion force stents having smaller feature widths and thicknesses in which manufacturing variations become proportionately more significant. In addition, a typical delivery catheter for use in expanding a stent includes a balloon folded into a compact shape for catheter insertion. The balloon is expanded by fluid pressure to unfold the balloon and deploy the stent. This process of unfolding the balloon causes uneven stresses to be applied to the stent during expansion of the balloon due to the folds causing the problem non-uniform stent expansion.
U.S. Pat. No. 5,545,210 discloses a thin-walled tubular stent geometrically similar to those discussed above, but constructed of a nickel-titanium shape memory alloy (“Nitinol”). This design permits the use of cylinders with thicker walls by making use of the lower yield stress and lower elastic modulus of martensitic phase Nitinol alloys. The expansion force required to expand a Nitinol stent is less than that of comparable thickness stainless steel stents of a conventional design. However, the “recoil” problem after expansion is significantly greater with Nitinol than with other materials. For example, recoil of a typical design Nitinol stent is about 9%. Nitinol is also more expensive, and more difficult to fabricate and machine than other stent materials, such as stainless steel.
All of the above stents share a critical design property: in each design, the features that undergo permanent deformation during stent expansion are prismatic, i.e., the cross sections of these features remain constant or change very gradually along their entire active length. To a first approximation, such features deform under transverse stress as simple beams with fixed or guided ends: essentially, the features act as a leaf springs. These leaf spring like structures are ideally suited to providing large amounts of elastic deformation before permanent deformation commences. This is exactly the opposite of ideal stent behavior. Further, the force required to deflect prismatic stent struts in the circumferential direction during stent expansion is proportional to the square of the width of the strut in the circumferential direction. Expansion forces thus increase rapidly with strut width in the above stent designs. Typical expansion pressures required to expand known stents are between about 5 and 10 atmospheres. These forces can cause substantial damage to tissue if misapplied.
FIG. 1 shows a typical prior art “expanding cage” stent design. The stent 10 includes a series of axial slots 12 formed in a cylindrical tube 14. Each axial row of slots 12 is displaced axially from the adjacent row by approximately half the slot length providing a staggered slot arrangement. The material between the slots 12 forms a network of axial struts 16 joined by short circumferential links 18. The cross section of each strut 16 remains constant or varies gradually along the entire length of the strut and thus the rectangular moment of inertia and the elastic and plastic section module of the cross section also remain constant or vary gradually along the length of the strut. Such a strut 16 is commonly referred to as a prismatic beam. Struts 16 in this type of design are typically 0.005 to 0.006 inches (0.127-0.1524 mm) wide in the circumferential direction. Strut thicknesses in the radial direction are typically about 0.0025 inches (0.0635 mm) or less to keep expansion forces within acceptable levels. However, most stent materials must be approximately 0.005 inches (0.127 mm) thick for good visibility on conventional fluoroscopic equipment. This high ratio of strut width to thickness, combined with the relatively high strut length and the initial curvature of the stent tubing combine to cause the instability and bucking often seen in this type of stent design. When expanded, the stent structure of FIG. 1 assumes the roughly diamond pattern commonly seen in expanded sheet metal.
Another stent described in PCT publication number WO 96/29028 uses struts with relatively weak portions of locally-reduced cross sections which on expansion of the stent act to concentrate deformation at these areas. However, as discussed above non-uniform expansion is even more of a problem when smaller feature widths and thicknesses are involved because manufacturing variations become proportionately more significant. The locally-reduced cross section portions described in this document are formed by pairs of circular holes. The shape of the locally-reduced cross section portions undesirably concentrates the plastic strain at the narrowest portion. This concentration of plastic strain without any provision for controlling the level of plastic strain makes the stent highly vulnerable to failure.
In view of the drawbacks of the prior art stents, it would be advantageous to be able to expand a stent with an expansion force at a low level independent of choice of stent materials, material thickness, or strut dimensions.
It would farther be advantageous to have a tissue-supporting device that permits a choice of material thickness that could be viewed easily on conventional fluoroscopic equipment for any material.
It would also be advantageous to have a tissue-supporting device that is inherently stable during expansion, thus eliminating buckling and twisting of structural features during stent deployment.
It would also be desirable to control strain to a desired level which takes advantage of work hardening without approaching a level of plastic strain at which failure may occur.
In addition, it would be advantageous to have a tissue-supporting device with minimal elastic recovery, or “recoil” of the device after expansion.
It would be advantageous to have a tissue supporting device that can be securely crimped to the delivery catheter without requiring special tools, techniques, or ancillary clamping features.
It would further be advantageous to have a tissue-supporting device that has improved resistance to compressive forces (improved crush strength) after expansion.
It would also be advantageous to have a tissue-supporting device that achieves all the above improvements with minimal foreshortening of the overall stent length during expansion.